Traditionally, analog radiographic imaging, e.g. in medical applications, is performed using a combination of a phosphor layer, converting the X-rays to light, and a photographic film, e.g. 35*43 cm for a relatively large field of view as used in chest x-rays. The light emitted by the phosphor is captured by the film which is developed to obtain an image on the film.
Direct methods using only a film without a phosphor layer exist but are characterised by a low efficiency and can not be used for medical applications due to the need of high radiation dose needed to expose the film.
Both methods have the drawback that the photographic film has to be chemically processed, leading to chemical waste products and loss of time.
The first digital X-ray systems, now known as computed radiography (CR), used a stored energy releasing phosphor sheet which is exposed to a radiation image during X-ray exposure. The stimulable phosphor stores the radiation image at exposure whereafter the stored energy image is read out using stimulating radiation scanning the phosphor plate, releasing the image-wise stored energy as light. The light is detected and an electronic image is generated by the light detector and processing electronics, whereafter it is digitised. There is always a loss of time between the X-ray exposure and the readout of the image. Changing of the phosphor plates for the subsequent exposures is also cumbersome and requires intervention by the operator of the imaging apparatus.
Direct digital X-ray systems (DR) reading out the radiation image in real time using electronic sensors do exist. The systems can be divided in two main classes:                1. Systems using direct X ray conversion to electrons. The incident radiation reaches a photoconductor material which transforms the radiation into electron-hole pairs in response to the intensity of e.g. the X-rays. Charges generated are collected in a silicon based chip supporting the photoconductor. Conversion materials typically used are amorphous Selenium, Mo, lead iodide, mercury iodide, CdTe CdZnTe etc.        2. Presently most systems use indirect conversion sensors having a scintillator or phosphor layer as conversion layer converting the X-ray into visible light which is then detected with an electronic light sensor array which converts the light into electrons. Used materials in the conversion layer may be e.g. Gadolinium oxysulfide or cesium iodide.        
It is essential that the conversion layer absorbs a significant part of the X-rays to achieve efficient detection, and thus the lowest possible dose to the patient. Well-designed electronic system should typically reach double efficiency versus film-based X-ray detection.
Detailed discussion of existing digital systems and drawbacks:
The majority of the current systems are flat panels made with amorphous Silicon technology. These are matrix arrays of a-Si addressing transistors covered with a photoconductor material in case of type 1. In case 2, these are matrix arrays of a-Si addressing transistors and a-Si photodiodes, covered with a phosphor. Size up to 43×43 cm is not a problem with these panels.                A first disadvantage however is that pixel resolution is currently limited, as amorphous silicon, due to its physical properties, does hardly allow to fabricate pixels smaller than some 100 microns square.        A second disadvantage lies in the fact that currently these panels are of the passive type, i.e. have no in-pixel amplification capability. This makes fast imaging (i.e. continuously producing more than 1 image per second) very cumbersome and expensive.        
In principle, a detector using crystalline Si or another high-quality semiconductor material would be far superior in terms of smaller pixel size reachable, and due to the possibility of in-pixel circuitry such as amplifiers which greatly helps for higher imaging speeds. The major problem is that the (silicon) wafers used to make such detectors are limited in size. Large wafers are very costly and also the yield of the production process can be low. I.e. the percentage of good sensors out of a production run decreases very rapidly with increasing size.                An existing solution, indeed making use of crystalline detectors such as CCDs or CMOS sensors for indirect detection, to this problem is that the light image generated by the conversion layer is reduced and imaged to the small image sensor by an optical system. This may be a lens system or an optical fibre system.        
The most important drawbacks of this method are:                Due to the reduction of the image a relatively low number of pixels are read out for a large image, resulting in an inferior image quality, due to lower resolution.        A significant portion of the light is lost, leading to lower detection efficiency and/or higher dose to the patient. This happens because with the current state of the technology, and with demagnifications higher than five to ten times, unless expensive cooling mechanisms are used, the electronic noise from the detector system will be so high in comparison with the electronic signal, that the intrinsic signal-to-noise ratio of the X-ray signal is significantly degraded.        The bulky optics does not allow that the readout system is housed inside e.g. a conventional X-ray cassette. This would be desirable however, as a system which can be housed in a cassette with the form factor of a conventional X-ray cassette could then be plugged into a conventional X-ray apparatus. This would allow an easy upgrade from analog to digital imaging workflow, whereby the existing X-ray apparatus does not have to be replaced entirely.        Standard cassettes are provided in several sizes for different applications and dimensions and specifications of cassettes are regulated by international approved standards such as ANSI Ph 1.49.        
In order to avoid undesirable low number of pixels mentioned above, systems have been sought to split up the large field of view into smaller sections, each covered by a separate imager. The method for doing this should however:                not show any visible greyscale response or differences between individual image sensors. This can usually be solved with proper calibration techniques.        Not lose any image information in between the image sensors. This is not so easy to do, and several approaches exist to try to circumvent this problem. These will now be described.        
Several prior art systems exist to make one large image using multiple image sensors:
Techniques try to minimise or even eliminate the spacing between the border pixels of the individual imagers by bringing them close together (butting).
Linear butting of sensors can be found in several documents:
In EP 262 267 an image-receiving plane is made up of a two-dimensional array of radiation detectors, each having his own image processing circuit. In between the sensors however insensitive gaps exist.
U.S. Pat. No. 6,207,944 describes a sensor in which the edge pixels are larger in order to provide butting of the different sensors. However by providing a larger edge pixel generates a distortion to the image at the location of the butting lines.
Likewise in U.S. Pat. No. 6,323,475 conductive tracks lead from the selected detector positions to offset readout circuit positions allowing for certain pixels to be bigger than others.
In WO 99/33 117 several imaging device tiles are very accurately mounted on a support structure.
Especially the very accurate mounting poses problems in fabrication and makes such a combination of imaging tiles costly. It is requires high-precision machining technology to cut the silicon very precisely. And high precision machinery is needed to mount the chips together. This is costly and complex.
In EPA-1 162 833 sensors are mounted to butt sensitive areas. In one direction the chips overlie each other while in the other direction the chips are butted. The problem is that typically the edges of chips can not be brought close together than 50-100 microns, leaving a gap in between that is not light-sensitive and thus could miss vital (in other words: possibly lethal, e.g. a small cancer trace) information.
Also other systems make use of overlap of sensors in one directions in order to bring pixels closer together and not to loose information. Such system can also be found in EP 421 869. A very accurate overlapping of the chips provides butting of the sensitive areas.
In U.S. Pat. No. 4,467,342 CCD chips for detecting radiant energy have an overlap joint without substantial phase difference occurring at the lap joint. The major difficulty is to align the sensors accurately. CCD sensors exhibit low efficiency for radiation image sensing.
Another approach is using multiple cameras providing overlapping images of a e.g. scintillator screen.
In U.S. Pat. No. 4,503,460, DE 20000603 and EP 618 719 a combination of plural television pickup devices coupled to one X-ray intensifier using optics is used. Various detected regions overlap. The system has the drawback that due to the use of lenses the apparatus is large and can not be retrofitted in existing machines and x-ray cassettes. These systems have the advantage that no information is missing in between the separate images. They have however also a lower light efficiency, the cost of the multiple systems and due to their thickness usually do not fit in a conventional film-based X ray system.
U.S. Pat. No. 6,038,286 makes use of mirrors to divide the image towards several camera systems. This system has nearly the same drawbacks as it uses multiple cameras. The height is however somewhat less. Overlap seam problems may however occur at the mirror's edges.
WO 91/10921 describes the use of the transparency of silicon when using DRAM Chips having large spaces between cells as imaging sensors. Large spaces in between cells is however not acceptable for medical applications.
Other systems use mechanical step and repeat systems successively positioning the sensors at different positions. Mechanical systems are however slow, expensive and pose on the long term reliability problems. These systems can e.g. not be used for in vivo diagnostics.
In conclusion, the current state of the art is that imagers are either butted and in such case there is always a compromise on the image detail at the seams, and positioning is cumbersome and expensive; or a camera-like approach is used, but in such case the detector assembly rapidly becomes too thick to allow insertion into conventional analog X-ray units.
There is thus a need for a digital x-ray detection system possible combining several desired properties:                having a relatively large sensing area with sufficient pixel density.        having low enough noise to produce more than one image per second,        being flat enough to allow insertion in conventional X-ray units,        not missing or significantly compromising parts of the image due to imperfect butting of the different imagers the detector plane is composed off, and        allowing at the same time low-cost manufacturing techniques for positioning the different imagers.        